Method and apparatus for assessing hemodynamic properties within the circulatory system of a living subject

ABSTRACT

An improved method and apparatus for non-invasively assessing one or more hemodynamic parameters associated with the circulatory system of a living organism. In one aspect, the invention comprises a method of measuring a hemodynamic parameter by measuring a non-calibrated value of the parameter non-invasively, and inducing a stress of the circulatory system while measuring a second parameter. The response of the circulatory system to the stress is determined directly from the subject, and a calibration function is derived from the response and applied to the non-calibrated measured value to produce a calibrated measure of the actual value of the hemodynamic parameter. Methods of generating a tissue transfer function from measurements of the subject, correcting for periodic or aperiodic error components, and providing treatment to the subject based on the measured hemodynamic parameters, are also disclosed.

BACKGROUND OF THE INVENTION

1. Field of the Invention

This invention relates generally to methods and apparatus for monitoringparameters associated with the circulatory system of a living subject,and specifically to the non-invasive monitoring of arterial bloodpressure.

2. Description of Related Technology

Arterial Blood Pressure Measurement

Several well known techniques have heretofore been used tonon-invasively monitor a subject's arterial blood pressure waveform,namely, auscultation, oscillometry, and tonometry. Both the auscultationand oscillometry techniques use a standard inflatable arm cuff thatoccludes the subject's brachial artery. The auscultatory techniquedetermines the subject's systolic and diastolic pressures by monitoringcertain Korotkoff sounds that occur as the cuff is slowly deflated. Theoscillometric technique, on the other hand, determines these pressures,as well as the subject's mean pressure, by measuring actual pressurechanges that occur in the cuff as the cuff is deflated. Both techniquesdetermine pressure values only intermittently, because of the need toalternately inflate and deflate the cuff, and they cannot replicate thesubject's actual blood pressure waveform. Thus, true continuous,beat-to-beat blood pressure monitoring cannot be achieved using thesetechniques.

Occlusive cuff instruments of the kind described briefly above havegenerally been somewhat effective in sensing long-term trends in asubject's blood pressure. However, such instruments generally have beenineffective in sensing short-term blood pressure variations, which areof critical importance in many medical applications, including surgery.

The technique of arterial tonometry is also well known in the medicalarts. According to the theory of arterial tonometry, the pressure in asuperficial artery with sufficient bony support, such as the radialartery, may be accurately recorded during an applanation sweep when thetransmural pressure equals zero. The term “applanation” refers to theprocess of varying the pressure applied to the artery. An applanationsweep refers to a time period during which pressure over the artery isvaried from overcompression to undercompression or vice versa. At theonset of a decreasing applanation sweep, the artery is overcompressedinto a “dog bone” shape, so that pressure pulses are not recorded. Atthe end of the sweep, the artery is undercompressed, so that minimumamplitude pressure pulses are recorded. Within the sweep, it is assumedthat an applanation occurs during which the arterial wall tension isparallel to the tonometer surface. Here, the arterial pressure isperpendicular to the surface and is the only stress detected by thetonometer sensor. At this pressure, it is assumed that the maximumpeak-to-peak amplitude (the “maximum pulsatile”) pressure obtainedcorresponds to zero transmural pressure. This theory is illustratedgraphically in FIG. 1. Note that in FIG. 1, bone or another rigid memberis assumed to lie under the artery.

One prior art device for implementing the tonometry technique includes arigid array of miniature pressure transducers that is applied againstthe tissue overlying a peripheral artery, e.g., the radial artery. Thetransducers each directly sense the mechanical forces in the underlyingsubject tissue, and each is sized to cover only a fraction of theunderlying artery. The array is urged against the tissue, to applanatethe underlying artery and thereby cause beat-to-beat pressure variationswithin the artery to be coupled through the tissue to at least some ofthe transducers. An array of different transducers is used to ensurethat at least one transducer is always over the artery, regardless ofarray position on the subject. This type of tonometer, however, issubject to several drawbacks. First, the array of discrete transducersgenerally is not anatomically compatible with the continuous contours ofthe subject's tissue overlying the artery being sensed. This hashistorically led to inaccuracies in the resulting transducer signals. Inaddition, in some cases, this incompatibility can cause tissue injuryand nerve damage and can restrict blood flow to distal tissue.

Other prior art techniques have sought to more accurately place a singletonometric sensor laterally above the artery, thereby more completelycoupling the sensor to the pressure variations within the artery.However, such systems may place the sensor at a location where it isgeometrically “centered” but not optimally positioned for signalcoupling, and further typically require comparatively frequentre-calibration or repositioning due to movement of the subject duringmeasurement.

Tonometry systems are also commonly quite sensitive to the orientationof the pressure transducer on the subject being monitored. Specifically,such systems show a degradation in accuracy when the angularrelationship between the transducer and the artery is varied from an“optimal” incidence angle. This is an important consideration, since notwo measurements are likely to have the device placed or maintained atprecisely the same angle with respect to the artery. Many of theforegoing approaches to lateral sensor positioning similarly suffer fromnot being able to maintain a constant angular relationship with theartery regardless of lateral position, due in many cases to positioningmechanisms which are not adapted to account for the anatomic features ofthe subject, such as curvature of the wrist surface.

Another significant drawback to arterial tonometry systems in general istheir inability to continuously monitor and adjust the level of arterialwall compression to an optimum level of zero transmural pressure.Generally, optimization of arterial wall compression has been achievedonly by periodic recalibration. This has required an interruption of thesubject monitoring function, which sometimes can occur during criticalperiods. This disability severely limits acceptance of tonometers in theclinical environment.

A further limitation of the tonometry approach relates to incompletepressure pulse transfer from the interior of the blood vessel to thepoint of measurement on the surface of the skin above the blood vessel.Specifically, even when the optimum level of arterial compression isachieved, there is incomplete and complex coupling of the arterial bloodpressure through the vessel wall and through the tissue, to the surfaceof the skin, such that the magnitude of pressure variations occurringwithin the blood vessel is different than that measured by a tonometricsensor (pressure transducer) placed on the skin. Hence, any pressuresignal or waveform measured at the skin necessarily differs from thetrue pressure within the artery. Modeling the physical response of thearterial wall, tissue, musculature, tendons, bone, skin of the wrist isno small feat, and inherently includes uncertainties and anomalies foreach separate individual. These uncertainties and anomalies introduceunpredictable error into any measurement of blood pressure made via atonometric sensor.

One prior art method of calibrating tonometric pressure measurementsutilizes an oscillometric device (i.e., a pressure cuff or similar) toperiodically obtain “actual” pressure information which is then used tocalibrate the tonometric measurements. This approach suffers from theneed to perform ongoing calibration events, specificallyinflations/deflations of the cuff, in order to maintain devicecalibration. Such calibration events are distracting, uncomfortable, andcan practically only be performed with a comparatively long periodicity.Furthermore, this technique does not calibrate based on measurement ofactual hemodynamic changes occurring within the circulatory system, butrather based on external measurements which may or may not berepresentative of the actual changes. No mechanism for correcting forincomplete pulse transfer from the blood vessel to the sensor(s) due tointerposed tissue, etc. is provided either.

Other prior art calibration techniques have attempted to transmit orinduce a perturbation within the blood flowing in the blood vessel, andsubsequently sense the component of that signal within the measuredhemodynamic parameter (e.g., blood pressure waveform) to generate anoffset or correction for the measured parameter. See, for example, U.S.Pat. No. 5,590,649 entitled “Apparatus and Method for Measuring anInduced Perturbation to Determine Blood Pressure” assigned to VitalInsite, Inc. ('649 patent). Under the approach of the '649 patent,changes in a variety of hemodynamic parameters resulting ostensibly fromchanges in blood pressure are modeled and stored within the device, andcompared to data obtained from a tonometric sensor. This approach,however, has a profound disability in that the calibration offset isdetermined not by direct measurement of the hemodynamic parameters ofthe subject under evaluation, but by modeling the relationship betweenblood pressure and perturbation wave velocity; i.e., velocity and phaseare modeled to have a certain relationship to changes in blood pressure;therefore, in theory, observed changes in velocity/phase of theperturbation wave can be used to generate estimations of actual bloodpressure within the subject being evaluated. The limits of this systemare clearly dictated by the ability to accurately model many complex,non-linear, interdependent parameters, as well as predict the timevariance of these many parameters.

Hemodynamics and Diseases of the Circulatory System

The science of hemodynamics, or the analysis of fluid (blood) flowwithin the body, is presently used effectively to detect and/or diagnosediseases of or defects within the circulatory system. For example,valvular disease, cardiac structural defects, venous disease, reducedcardiac function, and arterial disease may be assessed by studying howthe blood flows through various portions of the circulatory system. Ofparticular interest is the analysis of arterial diseases such asstenosis (i.e. blockage or reduction in effective cross-sectional areadue to arterial plaque, etc.). It is known that as the degree ofstenosis within the blood vessel of a living subject varies, certainchanges in the parameters of the circulatory system and in the overallhealth of the subject occur. As illustrated in FIG. 2, varying degreesof stenosis within a hypothetical blood vessel will occlude that bloodvessel to a generally proportional degree; i.e., no stenosis results inno occlusion and no attendant symptoms, while complete stenosis resultsin complete occlusion, with no flow of blood through the vessel and verydire symptoms in the subject. At levels of stenosis falling somewherethere between, the response can be somewhat more complex. For example,the subject may suffer stenosis which very significantly reduces theeffective cross-sectional area of a given blood vessel, yet manifestsitself in very few if any symptoms under normal levels of exercise.However, the same subject can exhibit dramatic symptoms with an increasein exercise. as the patient exerts more effort, the tissue underexertion has a higher metabolic demand requiring an increase inperfusion. Normally, vasodilation and collateralized blood flow providethe compensatory mechanism to increase the volumetric flow to meet thehigher volumetric demand. However, since the vessel is significantlystenosed, the compensatory mechanism has already been utilized to meetthe normal, non-exercise demand. As a result, the body is unable toincrease the volumetric demand since it has no way of minimizing theenergy loss associated with overcoming the resistance of the stenosed(decreased) area of the vessel. If volumetric flow does not increase,the increased metabolic demand is not met and the distal tissue becomesischemic.

By modeling the stenotic artery as a fluid system having an internalpressure (P) and blood mass flow rate (Q) or blood velocity (v), amodified version of the well known Bernoulli equation may be applied todescribe the flow of blood within the artery as follows:

ΔP∝4·ν²  Eqn. (1)

Hence, the foregoing relationship may be used to assess one hemodynamicparameter when another is known. For example, the pressure gradient (ΔP)across a stenosis within the artery may be estimated by obtaining dataon the velocity of blood flowing through the stenosis, and then usingthis velocity data within Eqn. (1). The velocity data may be obtained byany number of well known techniques, such as spectral Dopplerultrasound.

However, despite their utility in assessing the severity of stenosespresent in the artery and other such diseases, prior art hemodynamicevaluation techniques are effectively incapable of assessing theabsolute blood pressure within the artery at any given time. In theory,an accurate model of the response of the circulatory system could beused to estimate the value of parameters within the system (such as truearterial pressure) based on known or measured values of otherparameters. As can be appreciated, however, the circulatory system of aliving organism, and especially a human being, is extremely complex,with literally thousands of interconnected blood vessels. This systemincludes, inter alia, scores of capillaries, veins, and arteries, eachhaving their own unique physical properties. Furthermore, within each ofthe aforementioned categories of blood vessel, individual constituentsmay have markedly different properties and response within thecirculatory system. For example, two arteries within the human body may(i) have different diameters at different points along their length;(ii) supply more or less veins and capillaries than the other; (iii)have more or less elasticity; and (iv) have more or less stenosisassociated therewith.

The properties and response of each of the blood vessels also may beaffected differently by various internal and/or external stimuli, suchas the introduction of an anesthetic into the body. Even commonautonomic responses within the body such as respiration affect thepressure in the circulatory system, and therefore may need to beconsidered.

Considering these limitations, it becomes exceedingly difficult if notimpossible to accurately model the circulatory system of the human beingin terms of its fluid dynamic properties for use in blood pressureestimation. Even if a hypothetical circulatory system could beaccurately modeled, the application of such a model would be susceptibleto significant variability from subject to subject due to each subject'sparticular physical properties and responses. Hence, such approaches canat best only hope to form gross approximations of the behavior of thecirculatory system, and accordingly have heretofore proven ineffectiveat accurately determining the blood pressure within a living subject.

Based on the foregoing, what is needed is an improved method andapparatus for assessing hemodynamic parameters, including bloodpressure, within a living subject. Such method and apparatus wouldideally be non-invasive, would be continuously or near-continuouslyself-calibrating, and would be both useful and produce reliable resultsunder a variety of different subject physiological circumstances, suchas when the subject is both conscious and anesthetized. Lastly, suchimproved method and apparatus would be based primarily on parametersmeasured from each particular subject being assessed, thereby allowingfor calibration unique to each individual.

SUMMARY OF THE INVENTION

The present invention satisfies the aforementioned needs by an improvedmethod and apparatus for assessing hemodynamic properties, includingblood pressure, within a living subject.

In a first aspect of the invention, a method of assessing hemodynamicproperties including blood pressure within the circulatory system isdisclosed. The method generally comprises the steps of: measuring afirst parameter from the blood vessel of a subject; measuring a secondparameter from the blood vessel; deriving a calibration function basedon the second parameter; and correcting the first parameter using thederived calibration function. Once calibrated, the second parameter ismonitored continuously or periodically; changes in that parameter areused to indicate changes in the hemodynamic property of interest. In afirst exemplary embodiment, the first parameter comprises a pressurewaveform, and the second parameter comprises the total flow kineticenergy of blood within the blood vessel. During measurement of thepressure waveform, the blood vessel is applanated (compressed) so as toinduce changes in the hemodynamic properties within the blood vessel andcirculatory system; the kinetic energy during such applanation is thenmeasured and used to identify one or more artifacts within the pressurewaveform. A correction function is then generated based on theseartifacts, and applied to the measured pressure waveform to generate acorrected or calibrated waveform representative of the actual pressurewithin the blood vessel. In a second exemplary embodiment, the maximalvelocity of the blood flowing within the blood vessel is determinedusing an acoustic signal and used to derive a calibration function.

In a second aspect of the invention, an improved method of calibrating apressure signal obtained from a blood vessel of a living subject usingone or more measured parameters is disclosed. Generally, the methodcomprises: measuring a pressure waveform from the blood vessel;measuring a second parameter at least periodically from the bloodvessel; deriving a calibration function based on the second parameter;and correcting the first parameter using the derived calibrationfunction. In one exemplary embodiment, the method comprises measuringthe pressure waveform from a blood vessel of the subject; measuring asecond parameter from the same blood vessel at least once; identifyingat least one artifact within the pressure waveform based on the secondparameter; deriving a calibration function based on the measured secondparameter and at least one property associated with the at least oneartifact; applying the calibration function at least once to thepressure waveform to generate a calibrated representation of pressurewithin the blood vessel; and continuously monitoring the secondparameter to identify variations in blood pressure with time.

In a third aspect of the invention, an improved method of characterizingthe hemodynamic response of the circulatory system of a living subjectis disclosed. The method generally comprises the steps of: deriving afirst functional relationship between first and second parametersassociated with a blood vessel under certain conditions; measuring thefirst and second parameters non-invasively under those certainconditions; identifying at least one artifact within at least one of themeasured parameters; and scaling the measurement of the first parameterbased on at least the first functional relationship and the at least oneartifact.

In a fourth aspect of the invention, an improved method of calibrating ahemodynamic parametric measurement having an error component isdisclosed. Generally, the method comprises measuring a hemodynamicparameter associated with a blood vessel; identifying an error sourceassociated with the first parameter; generating a calibration functionbased on the error source; and correcting the measured hemodynamicparameter using the calibration function. In one exemplary embodiment,the method comprises measuring a pressure waveform from the bloodvessel; identifying a periodic variation associated with the kineticenergy (or maximal velocity) of the blood within the blood vessel overtime due to respiratory effects; generating a calibration function basedon synchronization with the variation in kinetic energy over time; andapplying the calibration function to the pressure waveform to correctthe waveform for the periodic variation. This respiratory effect is alsodetectable from the pressure signal, and potentially other signals aswell.

In a fifth aspect of the invention, an improved apparatus for measuringhemodynamic properties within the blood vessel of a living subject isdisclosed. The apparatus generally a first transducer for measuring afirst hemodynamic parameter associated with the blood vessel; a secondtransducer for measuring a second hemodynamic parameter associated withthe blood vessel; and a signal processor operatively connected to thefirst and second transducers for generating a calibration function basedon the signal produced by the second transducer, and applying thecorrection function to the signal produced by the first transducer. Inone exemplary embodiment, the blood vessel comprises the radial arteryof a human being, and the apparatus comprises a pressure transducerdisposed non-invasively in proximity thereto; an acoustic transduceralso disposed in proximity thereto; an applanation device used toapplanate the blood vessel; and a processor for processing signals fromthe pressure and acoustic transducers during applanation of the bloodvessel. The acoustic transducer transmits an acoustic emission into theblood vessel and receives an echo therefrom; information regarding theblood's velocity and/or kinetic energy during the applanation is derivedfrom the echo and used to generate a correction function which is thenapplied to the measured pressure waveform to calibrate the latter.

In a sixth aspect of the invention, an improved computer program forimplementing the aforementioned methods of hemodynamic assessment,modeling, and calibration is disclosed. In one exemplary embodiment, thecomputer program comprises an object code representation of a C⁺⁺ sourcecode listing, the object code representation is disposed on the storagedevice of a microcomputer system and is adapted to run on themicroprocessor of the microcomputer system. The computer program furthercomprises a graphical user interface (GUI) operatively coupled to thedisplay and input device of the microcomputer. One or more subroutinesor algorithms for implementing the hemodynamic assessment, modeling, andcalibration methodology described herein based on measured parametricdata provided to the microcomputer are included within the program. In asecond exemplary embodiment, the computer program comprises aninstruction set disposed within the storage device (such as the embeddedprogram memory) of a digital signal processor (DSP) associated with theforegoing hemodynamic measurement apparatus.

In an seventh aspect of the invention, an improved apparatus foranalyzing parametric data obtained according to the foregoing methodsand utilizing the aforementioned computer program is disclosed. In oneexemplary embodiment, the apparatus comprises a microcomputer having aprocessor, non-volatile storage device, random access memory, inputdevice, display device, and serial/parallel data ports operativelycoupled to one or more sensing devices. Data obtained from a subjectunder analysis is input to the microcomputer via the serial or paralleldata port; the object code representation of the computer program storedon the storage device is loaded into the random access memory of themicrocomputer and executed on the processor as required to analyze theinput data in conjunction with commands input by the user via the inputdevice.

In a eighth aspect of the invention, an improved method of providingtreatment to a subject using the aforementioned method is disclosed. Themethod generally comprises the steps of: selecting a blood vessel of thesubject useful for measuring pressure data; measuring the pressure dataof the subject non-invasively; generating a calibration function;applying the calibration function to the measured pressure data toproduce a calibrated representation of blood pressure within the bloodvessel; and providing treatment to the subject based on the calibratedestimate. In one exemplary embodiment, the blood vessel comprises theradial artery of the human being, and the method comprises measuring apressure waveform from the radial artery via a pressure transducer;using an acoustic wave to measure at least one hemodynamic parameter;deriving a calibration function based at least in part on the measuredhemodynamic parameter; calibrating the pressure waveform using thecalibration function to derive a calibrated representation of bloodpressure useful for diagnosing one or more medical conditions within thesubject; and providing a course of treatment to the subject based atleast in part on the calibrated representation.

BRIEF DESCRIPTION OF THE DRAWINGS

FIG. 1 is a composite graph illustrating the cross-sectional shape of anartery as a function of applied pressure and time according to the priorart “maximum pulsatile” theory.

FIG. 2 illustrates a blood vessel with varying levels of stenosis formedon the walls thereof.

FIGS. 3-3e are a logical flow diagrams illustrating one exemplaryembodiment of the method of assessing hemodynamic parameters within thecirculatory system of a living subject according to the invention.

FIGS. 4a and 4 b are graphs illustrating the relationship between bloodvelocity and reduction of the effective cross-sectional flow area of ablood vessel.

FIGS. 5a-5 c are graphs illustrating the relationship betweenapplanation pressure, cardiac sinus rhythm, and arterial walldisplacement according to the invention.

FIG. 5d is a graph illustrating the relationship between maximum bloodvelocity and percentage reduction in flow area (applanation pressure)for both diastolic and systolic pressures.

FIG. 6 is a cross-section of a portion of a typical human wristillustrating the relationship between the artery, skin, and interposedtissue and bodily components.

FIGS. 7a-7 b are graphs illustrating an exemplary transfer function andtransfer fraction, respectively, for the cross-section of FIG. 6, forboth systolic and diastolic conditions.

FIG. 8 is a graph of measured, actual, and scaled (calibrated) arterialpressure versus time for a typical human subject utilizing theinvention.

FIG. 9 is a logical flow diagram illustrating one exemplary embodimentof the method of modeling the hemodynamic response of the circulatorysystem of a living subject according to the invention.

FIG. 10 is a logical flow diagram illustrating one exemplary embodimentof the method of calibrating a hemodynamic parametric measurement forrespiration or other periodic error sources according to the invention.

FIG. 11 is plot of the velocity and kinetic energy (KE) of blood flowingwithin a typical blood vessel, illustrating the effects of respirationthereon.

FIG. 12 is a block diagram of one exemplary embodiment of the apparatusfor measuring hemodynamic properties within the blood vessel of a livingsubject according to the invention.

FIG. 13 is a functional block diagram of a second embodiment of theapparatus of FIG. 12 illustrating its use on the radial artery of ahuman being.

FIGS. 14a-14 b are perspective views of various machine readable mediahaving object code representations of computer programs incorporatingthe methods of the present invention.

FIG. 15 is a block diagram of a first embodiment of the apparatus foranalyzing parametric data according to the invention.

FIG. 16 is a logical flow diagram illustrating one exemplary embodimentof the method of providing treatment to a subject using theaforementioned methods.

DETAILED DESCRIPTION OF THE INVENTION

Reference is now made to the drawings wherein like numerals refer tolike parts throughout.

It is noted that while the invention is described herein in terms of amethod and apparatus for assessing the hemodynamic parameters of thecirculatory system via the radial artery (i.e., wrist) of a humansubject, the invention may also be embodied or adapted to monitor suchparameters at other locations on the human body, as well as monitoringthese parameters on other warm-blooded species. All such adaptations andalternate embodiments are considered to fall within the scope of theclaims appended hereto.

Overview

In one fundamental aspect, the present invention comprises a method ofassessing hemodynamic parameters within a living subject by artificiallyinducing “stresses” on the subject's circulatory system. The response ofthe circulatory system to these stresses is known or determinable, anduseful in identifying artifacts or markers with the observed data. Thesemarkers are subsequently used to calibrate measurements of theaforementioned hemodynamic parameters.

For example, as will be described in greater detail below, the presentinvention is useful at calibrating the blood pressure waveform obtainedfrom a tonometric or surface pressure sensor disposed over the radialartery of a human being, the non-calibrated pressure waveformpotentially varying substantially from that actually experienced withinthe radial artery itself. In one embodiment, the “stress” placed on theartery is applanation (i.e., compression), and the velocity of bloodflowing through the area of applanation is monitored to identify markerswithin the velocity profile. These markers correspond to, inter alia, astate of near zero transmural pressure across the artery wall. In thisfashion, an accurate measure of true arterial pressure may be obtainednon-invasively. It will be recognized, however, that the invention asdescribed herein may also be readily used in assessing other hemodynamicproperties, such as the pressure differential between two locationswithin a blood vessel, venous or arterial wall compliance, variations inthe strength of ventricular contraction, and the like, and accordinglyis not limited to the measurement of arterial blood pressure.

Method of Assessing Hemodynamic Properties

Referring now to FIG. 3, the method of assessing hemodynamic propertiesincluding blood pressure within the circulatory system according to theinvention is described. As shown in FIG. 3, the first step 302 of themethod 300 comprises measuring a first parameter from the blood vesselof a subject. In the present context, the parameter measured will be ablood pressure waveform derived from a pressure sensor or transducerdisposed in proximity to the radial artery of the subject, as describedin greater detail with respect to FIG. 3a herein. It will be recognized,however, that other hemodynamic parameters may be measured as previouslynoted. Implicit in the measurement of the first parameter is theexistence of one or more error sources; i.e., the measured value of theparameter is not wholly representative of, or differs from, the actualvalue of the parameter existing in the circulatory system of thesubject. In the instance of arterial blood pressure, the actual value isthat existing within the artery itself, as may be measured by the A-lineor “gold standard” technique of invasive arterial catheterization.Reasons for such errors or differences are discussed in more detailbelow with reference to FIG. 3a.

Next, in step 304 of FIG. 3, a stress is induced on the blood vesselwhich alters its hemodynamic properties (at least locally), therebyinducing changes in other parameters associated with the vessel orcirculatory system as a whole. As discussed with respect to FIG. 3b,this stress comprises in one embodiment applanating or variablycompressing the blood vessel as a function of time, thereby inducingchanges in, inter alia, the mass flow rate (Q), velocity (v) or velocitygradient, and kinetic energy (KE) of the blood in the region of theapplanation. It is noted, however, that stressors other than theapplanation stress previously described may be applied to the subject toaffect similar or other hemodynamic properties, such as, for example,circumferential occlusion (as would occur with a cuff-like device) toaffect arterial cross-sectional area, or the localized introduction ofchemical substances into the subject to affect the compliance of theartery. Many such stressor/hemodynamic parameter combinations may beused consistent with the invention.

Next, in step 306, a second parameter associated with the blood vesselis measured in order to facilitate derivation of a calibration functionin step 308 below. As discussed in greater detail with respect to FIG.3c herein, the second parameter in one embodiment comprises total bloodflow kinetic energy, since this parameter exhibits certain easilyidentified “artifacts” as a function of the application of the stressorin step 304. As used herein, the terms “artifact” and marker are usedsynonymously, and refer to any identifiable feature or relationshipexisting within a data set. Other parameters which exhibit the same orother artifacts may be used to derive the calibration function however,including, for example, maximum blood velocity, blood vesselcross-sectional area, and blood mass flow rate.

In step 308 of FIG. 3, a calibration metric or function is next derivedbased on the parametric information derived in step 306. Specifically,one or more artifacts or markers are identified within the parametricdata, these artifacts indicating when certain relationships between theactual and measured values of the first parameter of step 302 aboveexist. As will be discussed with reference to FIG. 3d herein, oneembodiment of the process of deriving a calibration function comprisesmeasuring total blood flow kinetic energy within the region of theapplied stressor (applanation), and identifying changes within thesystolic and/or diastolic velocity profiles as a function of theapplanation (correlated to percentage reduction of cross-sectional areaof the blood vessel).

In step 310 of the method of FIG. 3, the calibration function derived instep 308 is applied to the measurement of the first parameter of step302 to generate a corrected or calibrated measurement. Note that if thefirst parameter is measured continuously (or periodically) as a functionof time, the correction function of step 308 may be continuously orperiodically applied as appropriate, thereby generating a calibratedmeasurement of the first parameter in an ongoing or continuous fashion.However, due to a variety of different factors, both the actual “A-line”arterial pressure and the scale or magnitude of the required calibrationfunction may vary as a function of time; hence, any “calibrated”measurement based on the previously calculated calibration function willbe in error. In one alternative, the user may simply periodicallyrecalibrate by reapplying the stressor (e.g., performing anotherapplanation sweep), generating an updated correction function, andapplying this to the measured value of the first parameter.

However, as is described in greater detail herein below, the presentinvention advantageously provides the ability to generate a calibrationfunction at a first time t₁, and then monitor the second hemodynamicparameter (e.g., maximum velocity, kinetic energy, area, or flow)continuously for indications of variation of the measured parameter.This is accomplished in step 312 of the method 300 by controlling theexternal pressure applied to the artery so as to establish apredetermined relationship between true arterial and external pressure,as described further below.

In step 312, the pressure applied to the artery is controlled toselected value of the first parameter so as to maintain the pressureacross the artery wall (i.e., “transmural pressure”) within the arteryat or near a desired value. This process is referred to herein as“servoing” to a particular value. As discussed in detail with referenceto FIGS. 5a-5 d herein, this servoing generates a particular blood flowkinetic energy in the area of the applanation; changes in this kineticenergy are then used to identify changes in the true arterial pressure.This “continuous calibration” is a desirable attribute of the presentinvention, since the continued, accurate measurement of hemodynamicparameters with the blood vessel of a subject is of critical importance,especially in the context of surgery or other such life-threateningevolutions where arterial blood pressure is used as the basis formoment-to-moment decisions on treatment of the subject.

Referring now to FIG. 3a, one embodiment of the method of measuring oneor more hemodynamic parameters within a living subject (step 302 of FIG.3) is described. The first step 322 of the method 320 of FIG. 3acomprises selecting one or more hemodynamic parameters for measurement.Selection of the parameter(s) to be measured is a function not only ofthe condition to be assessed, such as the subject's blood pressure orseverity of stenosis with an artery, but also on the monitoring locationselected in step 324 below (i.e., certain parameters may only bemeasured at certain locations due to physical or other limitations, asin the case of a localized stenosis within an artery which is physicallylocated at a discrete point).

Next, in step 324, a blood vessel within the body of the subject isselected for monitoring. Due to its accessibility and relative proximityto the surface of the skin, the radial artery of the human being is anexcellent location for monitoring hemodynamic parameters within thecirculatory system, although it will be appreciated that other locationson the human being (or other species) may be used for this purpose. Asnoted above, the location of monitoring also may be related to ordetermined by the type of condition to be assessed or monitoring to beperformed. Of course, multiple monitoring locations may be employed,whether sequentially or in parallel, with the methods of the presentinvention.

With respect to the radial artery of the human being, it is furthernoted that anecdotal evidence suggests that the radial artery is onlyminimally affected by arterial diseases, including stenosis andcalcification due to diabetes. The reasons for this observed behaviorare beyond the scope of this discussion; however, this behavior is ofsome significance to the discussion of applanation stress providedherein with respect to FIG. 3b, since the presence of pre-existingarterial disease such as medial calcification could impact the abilityto accurately measure arterial blood pressure. By selecting the radialartery when performing blood pressure measurements, which utilizecontrolled applanation as the applied stress, the user is effectivelyinsulated from many potential error sources relating to pre-existingstenosis or calcification.

Next, in step 326 of FIG. 3a, one or more parametric sensors capable ofmeasuring or sensing the selected parameter(s) is/are disposed inproximity to the selected blood vessel. In the case of measuring bloodpressure on the radial artery of the human, a pressure sensor(transducer) is disposed physically in contact with the skin on theinterior surface of the wrist, so as to be atop the radial artery. Thetransducer may be one of the well understood piezoelectric type, or anyother type capable of producing a pressure signal in a known relation tothe pressure applied to the surface thereof. Methods and apparatus forpositioning the transducer(s) such that optimal signal coupling andsensing are achieved are also well known in the blood pressure measuringarts, and accordingly will not be described further herein. Note thatwhile in contact with the skin of the wrist, the transducer(s) areinitially maintained in a state of low or zero compression of theunderlying tissue/artery, for reasons to be more fully explained herein.

In step 328, a signal is measured from the transducer(s) as a functionof time. The signal may be measured discretely (e.g., at a predeterminedinterval) or continuously, depending on the desired frequency ofmonitoring. In the case of the exemplary pressure transducer previouslydescribed, the output signal for a continuous measurement will comprisea time variant waveform. In the case of arterial blood pressure, thewaveform will generally track the actual “gold standard” arterialpressure, yet will include error or offset which varies with thepressure changes according to the various phases of the cardiac cycle.This time variant, non-linear error, or “variable error” between themeasured and actual pressure waveform presents an additional complexityin the measurement process, one which the present invention isparticularly well adapted to overcome as will be described in greaterdetail below.

Referring now to FIG. 3b, one embodiment of the method of inducing oneor more stresses on the circulatory system of the subject (step 304 ofFIG. 3) is described in detail. In the first step 332 of the method 330,a stress to be applied is selected. As used herein, the term “stress”(or “stressor”) refers to any physical or physiological change withinthe circulatory system of the subject which is artificially induced. Inthe present embodiment, the stress to be applied comprises applanation,or physical compression of the selected (radial) artery as a function oftime. An applanation “sweep”, as used herein, generally refers to thesteady application of increasing or decreasing pressure to the artery ina direction generally normal to the surface of the skin overlying theartery. The concept of applanation is simply illustrated by one placingone wrist between the thumb and forefinger of the other hand, thumb atopthe interior portion of the wrist, and slowly increasing pressure on theradial artery until the artery is occluded. It will be recognized,however, that as a general proposition, applanation as used herein maytake on any variety of different forms, such as (i) a continuous linearrate of increasing or decreasing compression over time; (ii) acontinuous non-linear (e.g., logarithmic) increasing or decreasingcompression over time; (iii) a non-continuous or piece-wise continuouslinear or non-linear compression; (iv) alternating compression andrelaxation; (vi) sinusoidal or triangular waves functions; or (vi)random motion (such as a “random walk”). All such forms are consideredto be encompassed by the term “applanation.”

Referring to FIGS. 4a and 4 b, the hemodynamic effects of applanationare described in detail. As will be readily recognized, the increasingapplanation of an artery 400 such as the radial artery of the humanresults in a reduction in the effective cross-section of the artery.Similar to the arterial stenosis previously described, the applanation402 reduces the flow area within the artery, thereby resulting inincreased blood velocity (v) through the restriction to maintain aconstant volumetric flow. This relationship is well understood in thefluid dynamics art. The profile of velocity across the reducing flowarea is altered as well, as illustrated by the velocity gradient 404 ofFIG. 4a. Hence, a higher maximum velocity, a higher velocity gradient,and a greater energy or pressure gradient across the restricted flowarea result from applanation.

FIG. 4b illustrates the peak or maximum flow velocity within the arteryas a function of percent reduction of the flow area of the artery. Totalblood flow kinetic energy is similarly related to area due in part toits relationship to velocity, albeit somewhat more difficult to deriveas described in greater detail below. Note that for the purposes ofsimplicity in the present discussion, the percent reduction of flow areais assumed to be directly proportional to the applanation pressureapplied at the tissue (skin) surface, although in reality thisrelationship is substantially more complex as described further below.Further, FIG. 4b is generally illustrative of “steady state” operation,and does not examine the effects of variation in pressure due to, forexample, the normal cardiac cycle, also discussed in greater detailbelow.

As illustrated in FIG. 4b, in the region of low applanation pressure410, the percentage reduction of the flow area is small, and the effectson flow velocity and gradient are minimal. Volumetric blood flow (Q) isunaffected. As applanation pressure increases (region 412), the flowarea is further reduced, and while the volumetric flow is maintained,the blood velocity, velocity gradient, pressure gradient, and kineticenergy begin to increase correspondingly. As applanation pressurefurther increases, the flow area is substantially reduced, and velocity,velocity gradient, pressure gradient, and kinetic energy increasesubstantially, while still maintaining volumetric flow under normalmetabolic demand. In the stenotic artery, this region 414 corresponds to“subcritical” stenosis, i.e., the level of stenosis where the subject'sexcess volumetric capacity is significantly reduced, generally with fewor no attendant symptoms. The appellation of “sub-critical” refers tothe fact that the patient is asymptomatic with adequate tissue perfusionunder normal metabolic demand, and only becomes symptomatic when thedemand increases as occurs with exercise.

A further reductions in flow area produces a transition through what isknown as the “critical” region 416; in the critical region, the flowarea is so reduced so that there is inadequate energy to overcome theincreased flow resistance, and volumetric flow is no longer maintained.Between these regions 414, 416, a velocity “peak” 420 is formed.Anecdotal evidence suggests that this peak 420 occurs roughly at pointof 50% reduction in arterial diameter (corresponding roughly to 75%reduction in flow area). As a result, the blood velocity and thevolumetric flow, and the flow kinetic energy distal to the stenosed areadrop precipitously with further reduction in flow area. As the arterybecomes fully occluded and flow area approaches zero (region 418), thevolumetric flow Q approaches zero, as does blood velocity and flowkinetic energy.

Examination of FIG. 4b yields important information in terms ofcharacterizing one response of the circulatory system to one appliedstress. Specifically, the behavior of velocity as a function ofapplanation, and most notably the increase in maximum velocity withinthe velocity profile, allow the identification of the point where thepressure within the artery is effectively equal to that applied to thewall of the artery via external applanation. This condition is referredto herein as a condition of “zero transmural pressure”. During theapplanation sweep illustrated in FIGS. 4a and 4 b, a point is reached atwhich the external pressure applied to the exterior of the artery wallis just offset by the internal pressure within the artery. Until thispoint is reached, no significant reduction in flow area (and resultingattendant changes in velocity, velocity gradient, volumetric flow (Q),or kinetic energy as previously described) occurs. However, as theapplied pressure exceeds the arterial internal pressure, the diameterand cross-sectional area of the artery begin to be reduced, and themaximum flow velocity and velocity gradient begin to increase (region414 of FIG. 4b). This increase in maximum velocity (and kinetic energy)is used in the present embodiment as a “marker” of the point at whichthe transmural pressure is roughly equilibrated.

However, as previously discussed, the circulatory system is not a staticsystem, but rather dynamic and subject to significant intra-arterialpressure fluctuations, both due to the normal cardiac cycle, as well asother factors such as respiration (discussed below). Hence, suchpressure fluctuations must also be considered when measuring hemodynamicproperties, particularly intra-arterial pressure.

Referring now to FIGS. 5a-5 d, the response of the circulatory systemunder the aforementioned dynamic pressure fluctuations is described.FIG. 5a illustrates a normal sinus cardiac rhythm 500 for a human being.Within this sinus rhythm 500 are both systolic periods 502 and diastolicperiods 504 corresponding to various ventricular functions within theheart, as is well understood in the medical arts. These effectivelyrepresent maxima and minima within the sinus rhythm 500, and for theintra-arterial pressure.

FIG. 5b illustrates the displacement of the arterial wall as a functionof the aforementioned sinus rhythm 500 of FIG. 5a, and the externalapplanation pressure applied to the artery per FIG. 5c. Two opposedarterial walls 510, 512 are illustrated in FIG. 5b for sake of clarity,although they are effectively mirror images of one another in terms ofpressure response. As is well known in the art, arterial walls aretypically (in the healthy human) substantially compliant vessels havingsignificant elasticity and resiliency. Hence, as pressure within thevessel is increased, the opposing walls 510, 512 of the artery tend todeflect outward increasing the diameter of the artery, much as a balloonunder inflation. Similarly, as intra-arterial pressure is reduced, theresiliency of the artery walls reduces the diameter. It is well knownthat human arteries cyclically expand and contract to some degree duringthe normal cardiac cycle.

As shown in FIG. 5b, variations in blood pressure within the arterydeflect the walls of the artery outward to a maximum diameter 516corresponding to the systolic pressure 502, and allow the artery wallsto collapse to a minimum diameter 518 corresponding to the diastolicpressure 504. As the applanation pressure applied to the exterior of theartery (FIG. 5c) increases, the previously described condition of zerotransmural pressure is reached successively for both the systolic anddiastolic pressures. Specifically, with increasing applanation pressure,zero transmural pressure at the diastolic (lower pressure) condition 520is achieved first, followed by zero transmural pressure at the systolic(higher pressure) condition 522. Considering the diastolic condition 520first, as applanation pressure is increased beyond the zero transmuralpressure condition, the effective diameter (and flow area) of the arterybegins to progressively decrease, resulting in the increase in flowgradient and peak blood velocity and kinetic energy as previouslydescribed. As applanation pressure increases well above the diastolicpressure, the artery completely collapses during the diastolic portionof the cardiac cycle at point 527. Similarly, with increasingapplanation pressure, the diameter of the artery at the systoliccondition 522 also begins to decrease, with similar results, until theartery is completely collapsed under both diastolic portions 527 andsystolic portions 529 of the cardiac cycle. Based on the foregoingbehavior, two curves may be constructed (FIG. 5d) relating the variationin maximum blood velocity and percent flow area reduction (applanationpressure), both for the diastolic condition 520 and the systoliccondition 522. Note that the velocity “peak” 524 of the systoliccondition 522 occurs at a higher level of applanation than thecorresponding peak 526 for the diastolic condition 520, since greaterexternal pressure must be applied to collapse the artery in the formeras opposed to the latter. It is further noted that at pressures fallingbetween the systolic and diastolic maxima and minima of FIG. 5a, afamily of curves similar to those of FIG. 5d may be constructed, such afamily of curves being useful in characterizing the behavior of theartery and associated hemodynamic parameters during the entire cardiaccycle.

As with the velocity curve of FIG. 4b, the curves of FIG. 5d are usefulfor marking the point during the applanation sweep at which zerotransmural pressure is achieved, both during the diastolic and systolicportions of the cardiac cycle (or any portion there between). Theutility and application of this information is described in detail withreference to FIGS. 3d-3 e herein.

While the foregoing exemplary application of compressive or applanationstress is useful in the measurement of, inter alia, blood pressurewithin the selected artery, it will be recognized that other types ofstresses may be applied to induce response within the circulatorysystem. Artifacts or “markers” associated with these stresses may beutilized in a fashion generally analogous to that for the applanationstress; i.e., by correlating the presence of the markers or knownrelationships with certain hemodynamic conditions within the circulatorysystem in general or blood vessel in particular. Hence, the method ofFIG. 3b is in no way limited to the use of compressive stress.

Returning again to FIG. 3b, the second step 334 of the method 330 ofapplying stress to the selected blood vessel comprises providing amechanism by which such stress can be applied. In the context ofapplanation as described above, there is particular utility in using theaforementioned pressure transducer (used to measure the pressurewaveform) as the means by which the artery is applanated, since thisarrangement permits the pressure measurement to be made precisely at thepoint of applanation. An applanation mechanism of this type is describedherein with respect to FIG. 12. However, it will be appreciated that aseparate pressure transducer and applanation mechanism, or even otherconfigurations, may be used in conjunction with the present invention.

In step 336 of FIG. 3b, the provided mechanism is utilized to apply thestress to the selected artery. In the specific case of applanation, anapplanation “sweep” as previously described is applied, such that thepressure transducer is asserted at continually increasing levels ofpressure against the skin of the wrist, thereby compressing theunderlying artery. As with the method 320 of FIG. 3a, optimal placementand orientation of the applanation device over the artery may bedetermined using any variety of well understood prior art techniques,and accordingly is not discussed further herein. It is noted, however,that the foregoing method 320 maybe utilized even with non-optimaltransducer placement (e.g., by manual placement by the individualadministering treatment), so long as the signal coupling in such casesis adequate.

Referring now to FIG. 3c, one embodiment of the method of measuring asecond hemodynamic parameter associated with the blood vessel (step 306of FIG. 3) to facilitate derivation of a calibration function isdescribed. In the first step 342 of the method 340, the secondhemodynamic parameter to be measured is selected. Election of thisparameter is in some respects coupled to the selection of the firsthemodynamic parameter to be measured (FIG. 3a), as well as the selectionand application of stress on the circulatory system (FIG. 3b). In thecontext of blood pressure measurement and the use of compressive stress(applanation) as previously described, several “secondary” hemodynamicparameters may conceivably be used to generate a calibration function,including, without limitation, blood velocity, total blood flow kineticenergy, and blood volumetric flow rate, as well as any variations orcombinations thereof. Total blood flow kinetic energy is oneparticularly useful parameter to measure, as it contains one or morereadily observable markers of the zero transmural pressure condition orother useful relationships. The total flow kinetic energy is also lessprone to errors than certain other parameters, since it utilizesvelocity information obtained across the whole blood vessel, as well asthe amplitude information. Additionally, the peaking in the kineticenergy is more dramatic than the peaking in other parameters such asmaximum velocity.

Next, in step 344, the selected “secondary” parameter is measured usingan appropriate sensor or measurement technique. In the case of kineticenergy or blood velocity measurements, several well known techniquesexist to generally measure these parameters non-invasively. Ofparticular note is the use of acoustic energy (e.g., ultrasound) tomeasure blood velocity. Specifically, acoustic measurement techniquesgenerally employ the well known Doppler principle in measuring velocity,wherein the frequency shift associated with echoes reflected by theblood flowing within the blood vessel is analyzed to provide ameasurement of blood velocity. Numerous different variants of acousticblood velocity measurement techniques exist, including the use of acontinuous acoustic wave (CW), and acoustic pulses (pulsed Doppler).Such techniques are well known and understood, and accordingly will notbe described further here.

Similarly, acoustic measurement techniques may be used to derive ameasurement of the kinetic energy of the blood flowing within thesubject blood vessel. It is noted that as a result of the complex bloodvelocity gradient created with in the blood vessel during applanation(FIG. 4a), calculation of the kinetic energy of the blood within theblood vessel as a whole is not simply proportional to the square of themaximum blood velocity described above; rather, estimation of thekinetic energy requires the application of summation or integrationtechniques which capture the complexity of this gradient. Suchsummation/integration techniques for calculating blood kinetic energyare well known in the art, and accordingly are not described furtherherein.

In another embodiment, the applanation (external) pressure at which thedesired marker is exhibited may be determined using time-frequencymethodology as described in Applicant's co-pending U.S. patentapplication, Ser. No. 09/342,549, entitled “Method And Apparatus For TheNoninvasive Determination Of Arterial Blood Pressure” filed Jun. 29,1999, and incorporated herein by reference in its entirety. Using thistime-frequency methodology, the applanation pressure at which thetransmural pressure equals zero can be determined by constructingtime-frequency representations of the acoustic energy reflected withinthe artery. When the time-frequency distribution is maximized, the zerotransmural pressure condition is achieved. Hence, the maximaltime-frequency distribution acts as yet another marker for the purposesof the present invention.

In yet another embodiment, the so-called acoustic “A-mode” may be usedto monitor the second hemodynamic parameter. In this approach, acousticwaves are generated and transmitted into the blood vessel; reflectionsor echoes from the transmissions are received and analyzed to determinethe relationship between the time of transmission and the time ofreceipt. Through such analysis, the relative diameter of the artery atdifferent points in time, and different points within the cardiac cycle,can be determined. Analogous to the well known time domain reflectometer(TDR), the A-mode technique utilizes reflections generated by thetransition of an acoustic wave across various boundaries betweenmaterials of different acoustic properties (e.g., the “near” arterywall/tissue boundary, the “near” artery wall/blood stream boundary, theblood stream/“far” artery wall boundary, etc.). Specifically, therelative timing of these reflections is analyzed to determine thedistance between the various boundaries. Knowing the propagation speedof the acoustic wave through the different media, the distance betweenthe reflective boundaries (i.e., tissue thickness, artery diameter,etc.) can be determined. Recalling that per FIG. 5b, the deflection ofthe artery walls (under both systolic and diastolic portions of thecardiac cycle) varies as a function of applanation pressure, changes inthe arterial diameter (and area, related thereto) may be used as“markers” of the zero transmural pressure condition, or other conditionsof significance, analogous to the use of increasing maximum velocity toidentify such conditions. Specifically, when the diameter of the arteryjust begins to decrease, the externally applied pressure just slightlyexceeds the internal arterial pressure at that point in time.

It will further be recognized that other acoustic modalities may beemployed in conjunction with the invention described herein, includingfor “M-mode” (motion mode) or “B-mode” (brightness mode) both of whichare well known in the acoustic signal arts.

Despite the use of acoustic waves in each of the foregoing embodimentsfor measuring the secondary hemodynamic parameter and markers associatedtherewith, it will be recognized that other non-acoustic techniques maybe applied to identify such markers. For example, other methods ofaccurately measuring arterial diameter/area, such as usinginterferometry, may be employed to identify the zero transmural pressurecondition. All such techniques are considered to fall within the scopeof the present invention.

Referring now to FIG. 3d, the stressor magnitude at which the desiredhemodynamic condition is achieved (e.g., applanation pressure at whichzero transmural pressure is achieved) is correlated to the actual ortrue arterial pressure. In the simple case where there is a high degreeof coupling between the applied stress and the stress actually felt bythe blood vessel, the measured stress can be equated to the actualstress. Specifically, in the context of arterial blood pressuremeasurements where applanation (compressive) stress is applied, thepressure applied by the applanation device and sensed by the associatedpressure transducer could be equated to the actual arterial pressurewhen the artifact or “marker” condition is observed. For example, ifincreasing blood kinetic energy correlates to a condition of zero ornear-zero transmural pressure as previously discussed, the pressureapplied against the artery wall when such increase in kinetic energy wasobserved would equate to true intra-arterial pressure. Hence, if thecoupling between the point of pressure application (e.g., skin) and theartery wall was very high, the pressure applied at the point ofapplication would approximate that applied to the artery wall, andtherefore would also approximate the pressure within the artery.

However, as previously discussed, the tissue, tendons, and skininterposed between the artery wall and the pressure transducer in manycases create a complex relationship between the pressure applied by thetransducer (or applanation mechanism) and the pressure actually felt bythe artery wall. Simply stated, some of the pressure applied to the skinis used to compress this interposed material; hence, only a portion ofthe externally applied pressure is actually felt by the artery wall.Additionally, it is noted that tissue is also present below the bloodvessel and above bone; some loss occurs in compressing this tissue aswell.

Therefore, depending on the tissue compliance and degree of coupling fora given subject, a certain amount of error in the measurement ofarterial pressure will be introduced when basing such a measurement onthe externally applied pressure (e.g., that measured by the pressuretransducer).

One prior art approach to this problem was to model the response ofinterposed material (for example, as a system of springs having linearforce constants), and correct the pressure measured by the pressuretransducer based on this model. This approach, however, is only as goodas the model used; different subjects with different tissue thickness,density, and compliance values (as well as the location of the tendonsand bone relative to each other and the artery) will responddifferently, and these differences are not accounted for in such models.Furthermore, even for a single subject, changes in the response of thetissue and arteries of that subject may occur over time or as a functionof externally induced stresses. For example, when an anesthetic isintroduced into the circulatory system of the subject, a given arterymay become substantially more compliant, thereby losing much of itsresiliency. This change in compliance alters the relationship betweenactual and measured arterial pressure, and accordingly reduces theaccuracy of any blood pressure estimate based thereon.

In contrast, the methodology of the present invention overcomes thissignificant limitation by measuring the actual response of theinterposed tissue and material for each subject as opposed togenerically modeling it as in the prior art. Specifically, the presentinvention generates a functional representation of tissue and arterialcompliance based on actual compression of these components.

In the exemplary embodiment of the method 350 illustrated in FIG. 3d,the aforementioned “A-mode” acoustic transmission is used to monitor thecompression of each of the components interposed between the applanationdevice and the artery interior wall. The compression of these components(step 352) proceeds generally according to their individual materialproperties, which are unknown and interdependent and thereforeexceedingly complex to model. However, by making direct observations ofthe actual compression of these components, the transfer functionexisting between the externally applied force and the force applied atthe interface of the artery wall and the pressurized fluid (blood)within the artery can be approximately determined for each individual,and for the specific location being applanated. As illustrated in FIG.6, the region between the interior wall of the artery and the surface ofthe skin above the artery may be divided into several discrete regions,such as the skin 602, tissue 604, and artery wall 606. The distances d₁,d₂, d₃ and d₄ between the surface of the skin 608 and the skin/tissueboundary 610, the skin/tissue boundary and the tissue/artery boundary612, the tissue/artery boundary and the artery/blood boundary 614, andthe artery/blood boundary 614 and the blood/artery boundary 616,respectively, are measured in step 354 using A-mode acoustictransmissions which identify reflections from these boundaries, aspreviously described. Additionally, the relative location of bone 620and tendon (not shown) have great influence on the transfer loss. Ineffect, a restoring spring force of sorts exists between the tendon andtissue 622 and bone 620 and tissue 622. The loss of pressure transfer isat least partially associated with overcoming these restoring forces, aswell as with the compliance of the tissue. Hence, during applanation(and during specific portions of the cardiac cycle), a transfer functionbetween artery diameter (and flow area) and applied external pressure isdeveloped per step 356. Specifically, for the diastolic and systolicportions of the cardiac cycle, different transfer functions will existas illustrated in FIGS. 7a-7 b. At low applied pressure (as measuredrelative to the actual intra-arterial pressure), relatively littlecompression of interposed tissue, underlying tissue/tendon, artery wall,etc. has occurred, and hence further increases in applied pressuregenerally contribute disproportionately to further compression of thesecomponents. At higher values of applied pressure, the interposedcomponents are substantially compressed, and a relatively small fractionof any further increases in applanation pressure is used to compress theinterposed and underlying components. Hence, in general, the “transferfraction”, or the ratio of transferred pressure to applied pressure,increases as a function of applied pressure, as illustrated in FIG. 7b.In the theoretical case of free-floating incompressible materialsinterposed between the pressurized blood in the artery and thetransducer, the transfer fraction would be 1:1, indicating completecoupling.

The foregoing derived transfer function, can then be utilized to correctthe error of the incomplete pressure transfer measured by the pressuresensing introduced by the interposed tissue, etc., by identifying theregions of interest per step 358. For example, if the zero transmuralpressure condition within the artery during the diastolic portion of thecardiac cycle is achieved when a pressure of 60 mm Hg is measured, thetrue diastolic pressure will be some percentage higher, where thepercentage is determined by the degree of pressure transfer loss. Thetransfer fraction for that monitoring location indicates the fraction orpercentage of the intravascular pressure which is transferred to thesurface of the pressure measuring sensor.

Note that the transfer function and/or transfer fractions may berepresented and stored in any variety of different formats aftermeasurement, such as in look-up tables in a digital random access memoryas described further below with respect to the apparatus of FIG. 12.Furthermore, it will be readily appreciated that while the method 350described above is used to determine the transfer fraction for one ormore discrete pressure conditions (i.e., systolic and/or diastolicpressures), the transfer fraction may be readily determined for a rangeof pressures, thereby forming a transfer function as a function ofpressure, as described in greater detail below. Hence, if the bloodpressure of the subject does vary, the present invention utilizes thistransfer function to correct the measured value of pressure within anypressure range.

Similarly, it will be recognized that methods of determining thetransfer function/fraction other than the A-mode acoustic technique maybe utilized, either alone or in conjunction with the A-mode technique.

In sum, the method 350 of FIG. 3d involves determining a transferfunction/fraction as related to applied stress (e.g., pressure) for thesubject and location being monitored, and calibrating the measuredparameter at the designated “marker” point using the transferfunction/fraction to determine the actual value of the parameter. In thecase of blood pressure monitoring, this process involves applying anapplanation pressure at which the kinetic energy term begins to increase(or alternatively, the maximum blood velocity begins to increase, theflow area begins to decrease, or some other desired condition isobserved), and then correcting the measured value of the measuredpressure using the transfer fraction to determine the actualintra-arterial pressure. When considered over the entire cardiac cycle,this method 350 produces a scaling or “stretching” function which isapplied to the entire measured pressure waveform 800 to calibrate it tothe true intra-arterial pressure 801, and thereby produce a “calibrated”waveform 806 as shown in FIG. 8. It is noted that depending on theportion of the measured pressure waveform being considered (e.g.,diastolic portion, systolic portion, or there between), the ratio ofactual or A-line intra-arterial pressure to the measured pressure willvary. This concept is graphically illustrated in FIG. 8, wherein theratio of amplitudes at the systolic portion of the cardiac cycle R₁ 802is not equal to the ratio of amplitudes at the diastolic portion R₂ 804.

It should be noted that while certain circumstances and individualsubjects require the determination and application of a transferfunction as described with respect to FIG. 3d, the general methodologyof the invention may potentially be applied in some cases without atransfer function. For example, where a subject has a high degree ofcoupling between the skin and artery wall, the error associated with thepressure measured via the transducer placed at the skin surface may onlyconstitute a small fraction of the total measurement, and wouldtherefore be acceptable in certain monitoring environments. Hence,calculation and application of the transfer function is not arequirement of the present invention under all circumstances.

Referring now to FIG. 3e, the method of continuously calibrating thehemodynamic parameter being measured is described. As discussed withreference to FIG. 3d above, the transfer function is useful forcorrecting the measured pressure waveform for compression of theinterposed tissue, artery wall, etc. This transfer function is obtainedduring an applanation sweep performed at a given monitoring location onthe subject, such as the radial artery. However, to permit continuousmonitoring of the subject's arterial blood pressure, a mechanism isneeded whereby changes in the measured parameter can be accuratelyobserved and scaled between calibration events (e.g., applanationsweeps).

As previously discussed, prior art calibration approaches relied onperiodic calibration events (such as asculatory cuff measurements) to“continuously” calibrate the measured pressure waveform. Theterm“continuously” used with reference to these systems is somewhat of amisnomer, since what actually occurs is periodic (rather thancontinuous) updates of the scaling function. This approach presents atleast one serious defect, that being the lack of calibration during theinterval between periodic calibration updates. Depending on theactivities of the subject being monitored, their true arterial bloodpressure may vary significantly in a short period of time, and in somecases in a rapid or prompt fashion. For example, during surgery, actionsby the surgeon such as artery re-section may have profound effects onthe circulatory system of the subject, including their arterial bloodpressure. Similarly, the difference between pre-induction (i.e.,pre-anesthesia) and post-induction blood pressure values may bedramatically different, due in large part to the change of compliancewithin many of the arteries in the subject's body resulting from theanesthetic.

Since the prior art approaches in no way monitor the actual hemodynamicproperties occurring within the artery, if such significant changes intrue arterial blood pressure occur between periodic calibration events,they in many cases will go undetected. Rather, such prior art approachestypically monitor blood pressure tonometrically, these measurementsbeing potentially very different from true arterial pressure. The priorart systems typically adjust the scaling factor or calibration toaccount for the measured change in tonometric blood pressure (which mayor may not be close to true blood pressure). The result of this methodis to produce so-called “calibrated” blood pressure values which in factare not calibrated, but comprise a widely varying scaling component.This failure to track actual or true arterial blood pressure betweencalibration events can be catastrophic in cases where minute-to-minutemeasurements of blood pressure may be critical, such as during surgery.

The methodology of the present invention overcomes the foregoingsignificant limitations of the prior art by using the measured“secondary” hemodynamic parameter previously described to track changesin the first or “primary” measured hemodynamic parameter (e.g., bloodpressure), as described in detail below.

In one embodiment, the kinetic energy of the blood is monitored usingthe aforementioned acoustic (or other) techniques while the zerotransmural pressure state (or some other state determined to be ofsignificance) is maintained within the artery, as illustrated by themethod 370 of FIG. 3e. Specifically, the applanation device, which inthe embodiment described below with respect.to FIG. 12 also comprisesthe pressure and ultrasonic transducers, is “servoed” or continuallymodulated against the skin above the monitored artery in step 372 so asto maintain the desired pressure condition. The measured(non-calibrated) primary parameter, here pressure, is monitored as afunction of time at the same time per step 374. Depending on theparticular application, the modulation of step 372 may be controlled soas to maintain the transmural pressure at a specific value during thediastolic portion of the cardiac cycle, or alternatively during thesystolic portion of the cycle. As yet another alternative, theapplanation device may be modulated or servoed to maintain the meantransmural pressure (calculated over one or more complete cardiaccycles) at a predetermined value. Servoing may also be conducted tomaintain a desired maximal blood velocity condition, or cross-sectionalarea condition. Many other such “target” servo values may be substitutedwith equal success, and the choice of such values, as well as theparametric relationship on which this value is based. (e.g., the regionon the maximum velocity v. flow area plots of FIG. 5d in which it isdesired to operate) is solely determined by the needs of the user andthe particular application in which the method is employed.

Next, in step 376, the secondary hemodynamic parameter is measured as afunction of time using a suitable technique. In the present embodiment,the total kinetic energy (or maximum blood velocity) is measured usingan acoustic Doppler system of the type previously described.

In step 378, the value of the secondary parameter measured in step 376is analyzed to identify changes in the primary parameter. For example,when the applanation device is servoed to maintain zero transmuralpressure in the diastolic portion of the cardiac cycle, changes inkinetic energy are used to track changes in intra-arterial bloodpressure. The results of this analysis are compared to predeterminedacceptance or control criteria per step 380 to determine if furtheradjustment of the applanation device is required (step 382). Forexample, if significant increases or rates of increase in total bloodflow kinetic energy were observed in steps 378-382 (thereby indicatingthat the applanation pressure felt by the artery wall was exceeding thetrue intra-arterial pressure), then the applanation pressure could bereduced so as to maintain the artery at a near-zero transmural pressurecondition, as reflected by smaller increases or rates of increase inkinetic energy. It will be recognized that any type of control schemewhich controls one parameter based on measurements of one or more otherparameters may be used to effect the desired behavior, including fuzzylogic or PID controllers of the type well known in the control systemarts.

Notwithstanding the foregoing, it will be recognized that the continuouscalibration of the first hemodynamic parameter using the method of FIG.3e may be accomplished using additional or other secondary parametersincluding, for example, maximal blood velocity and/or arterialcross-sectional (flow) area.

It is also again noted that in contrast to prior art approaches, thetechniques of FIGS. 3-3e discussed above advantageously involve nomodeling or estimation of parameters within the circulatory system ofthe subject being monitored; all information is derived via directmeasurement of the subject at the selected monitoring location, andtherefore is particularly adapted to that individual and that location.

Method of Characterizing Hemodynamic Response of Circulatory System

Referring now to FIG. 9 a method of characterizing the hemodynamicresponse of the circulatory system of a living subject is disclosed. Asillustrated in FIG. 9, the first step 902 of the method 900 comprisesderiving a first functional relationship between first and secondparameters associated with a blood vessel in relation to an appliedstress. In the context of arterial blood pressure measurement, the firstfunctional relationship derived in step 902 comprises therelationship(s) between arterial cross-sectional area (applanationpressure) and total blood flow kinetic energy as previously describedherein, although it will be recognized that any number of differentfunctional relationships may be substituted therefor. For example, thefunctional relationship between maximal blood velocity and flow area,velocity gradient and flow area, or volumetric blood flow (Q) and flowarea, may be used if desired.

Next, in step 904, one or more artifacts or markers present within thefunctional relationship derived in step 902 above are identified. In thecase of arterial blood pressure measurement as previously described, theartifact comprises the increasing kinetic energy or blood velocity afterthe condition of zero transmural pressure is achieved for the diastolicand/or systolic conditions. These artifacts comprise points for thecalibration function previously described with respect to FIGS. 3c-3 eherein. Alternatively, the points at which wall diameter begins todecrease at the systolic and diastolic portions of the cardiac cycle, asmeasured by A-mode ultrasound or other similar techniques, mayconstitute a marker of zero transmural pressure.

Next in step 906, one of the functionally related parameters from step902 above is measured non-invasively as a function of the stressapplied. In the above-referenced example, this measurement wouldcomprise measuring blood velocity within the artery as a function oftime (and applanation pressure), and deriving total flow kinetic energytherefrom.

Lastly, in step 908, the calibration “function” (which in theory may beas few as one data point) is applied to the measured response of aselected parameter associated with the circulatory system based on theartifact identified in step 904, thereby producing a calibratedcharacterization of the response of that parameter. For blood pressure,the selected parameter is tonometrically measured (i.e., non-calibrated)pressure, and the calibrated characterization comprises calibrated (or“true”) arterial blood pressure determined at, inter alia, the pointwhere the kinetic energy of the blood begins to increase.

Furthermore, the effects of potential errors (such as that due toincomplete signal transfer due to tissue compliance) may be accountedfor as part of step 908 as well.

Method of Calibrating for Periodic Error Sources, Including Respiration

Referring now to FIG. 10, a method of calibrating a hemodynamicparametric measurement for periodic error sources is disclosed. Thefirst step 1002 of the method 1000 comprises measuring a firsthemodynamic parameter associated with a blood vessel. As previouslydescribed, this parameter may comprise arterial blood pressure, oranother parameter such as differential pressure, etc. In the case ofarterial blood pressure, this parameter is the non-calibrated pressurewaveform measured using the tonometric pressure transducer.

Next, in step 1004, a second hemodynamic parameter is measured on thesubject, as previously described. This second hemodynamic parameter maycomprise kinetic energy, maximum blood velocity, arterial diameter, flowarea, etc. In one embodiment, the kinetic energy is calculated based onmeasurements of blood velocity made using Doppler ultrasound.

Next, in step 1006, periodic error sources associated with the firstparameter are identified within the second parameter. In one exemplarycase, the periodic error source relates to the respiration of thesubject being monitored, illustrated in FIG. 11. As shown in FIG. 11,the velocity and kinetic energy of the blood flowing within the radialartery of a human being generally includes a time-variant, periodiccomponent. This periodic behavior is due in substantial part to therespiration cycle of the subject, and occurs at much lower frequencythan the typical cardiac cycle. Hence, the normal cardiac cycle 1102 canbe thought to be “amplitude modulated” by the periodic respiratoryvariance 1104.

The origin of the respiratory periodic variance relates to the varyingpressures which occurs as the diaphragm ascends and descends. Withinspiration, the diaphragm should descend, increasing intra-abdominalpressure and decreasing intra-thoracic pressure. The increase in thepressure differential from the abdomen to the right atrium increases thevolumetric flow back to the right atrium. With expiration, as thediaphragm ascends, the intra-abdominal pressure decreases and theintra-thoracic pressure increases. The result is more venous return tothe abdomen from the lower extremities, but less return to the rightatrium. The cyclical changes in volume and pressure are reflectedeverywhere throughout the circulatory system, since it is a closedsystem.

The aforementioned cyclical respiratory changes result in variant flowvelocities and kinetic energies for, inter alia, the measured diastolicand systolic pressures. In a normal adult human being, anecdotalevidence obtained by the Applicant herein suggests that the magnitude ofsuch variations may be on the order of 20 mm Hg or more in severe cases.Taken as a fraction or percentage of the systolic and diastolicpressures, this variation in pressure due to respiration may besignificant, especially for the lower diastolic pressures measured whenthe subject is not ambulatory, such as during surgery.

These variations are accounted for in the present invention, whenrequired, by synchronizing the derivation of the calibration functionfrom the measurement of the secondary hemodynamic parameter (e.g.,velocity, kinetic energy, or area). Specifically, in step 1008 of themethod 1000, the periodicity of the respiratory variation is analyzedand determined, and this information is used to synchronize thederivation of the calibration function to a common point on the period(“carrier”) respiration waveform. Identification of the respiratorycomponent and its periodicity is accomplished using any one of a numberof algorithms well known in the signal processing arts; accordingly,such algorithms will not be discussed further herein. It is noted thatsince the respiratory rate and/or “depth” of respiration of the subjectmay vary with time, thereby affecting the periodicity and magnitude ofpressure/flow variations within the artery, the periodicity of therespiratory effect should be continually (or at least frequently)calculated.

Next, in step 1010, a calibration function is developed based onmeasurements of the secondary hemodynamic parameter taken at theperiodicity prescribed by the result of step 1008. For example, a seriesof blood velocity measurements may be taken every 7 seconds (eachmeasurement corresponding to the same relative point on the respirationwaveform, but displaced in time), and this information used to derivekinetic energy values and a calibration or “stretching” function asdescribed previously herein with reference to FIGS. 3-3e.

Lastly, in step 1012, the stretching function of step 1010 is applied tothe measured (i.e., non-calibrated) waveform of step 1002. Note that byvirtue of measuring the second hemodynamic parameter at a similar pointrelative to the respiration waveform, the effects of respiration acrossthe entire respiration cycle are accounted for. Hence, the derivedstretching function may be applied to the entire non-calibrated pressurewaveform), as opposed to only those portions of the waveformcorresponding to the points in time when the second parameter wasactually measured. Assuming the pressure transfer to be relativelylinear around the systolic pressure variations with respiration, and thediastolic pressure variations with respiration, no other correctionwould be necessary. An additional correction can be applied if thenon-linearities are significant enough, by calculating the correctionfactors at a different phase of the respiration cycle. This represents asignificant advantage in providing a continuously (as opposed toperiodically) calibrated representation of true arterial blood pressure.

It will be appreciated that while the foregoing discussion is cast interms of periodic error due to respiratory system effects, other typesof errors, periodic or aperiodic, may be accounted for using themethodology of the present invention as illustrated in FIG. 10. Forexample, the effects of an arrhythmia within the heart of the subjectmay be identified and accounted for during derivation of the calibrationfunction. An arrhythmia within the heart of the subject may beidentified using signal processing algorithms specifically adapted forthe purpose of identifying aperiodic components within waveforms, suchalgorithms being well known to those of ordinary skill in the signalprocessing arts. Numerous other. types on non-periodic error componentsmay also be identified in conjunction with the method of FIG. 10.

Apparatus for Hemodynamic Assessment

Referring now to FIG. 12, an apparatus for measuring hemodynamicproperties within the blood vessel of a living subject is described. Inthe illustrated embodiment, the apparatus is adapted for the measurementof blood pressure within the radial artery of a human being, although itwill be recognized that other hemodynamic parameters, monitoring sites,and even types of living organism may be utilized in conjunction withthe invention in its broadest sense. The apparatus 1200 of FIG. 12fundamentally comprises a pressure transducer 1202 for measuring bloodpressure from the radial artery tonometrically; an applanation device1204 coupled to the transducer 1202 for varying the degree ofapplanation (compression) on the artery; an acoustic transducer 1206 forgenerating acoustic emissions and reflections thereof, these acousticemissions being used to derive blood velocity (and kinetic energy); asignal processor 1208 operatively connected to the pressure and acoustictransducers 1202, 1206 for analyzing the signals generated by thesetransducers and generating a calibration function based thereon; asignal generator/receiver 1210 used to generate acoustic signals fortransmission into the artery, and receive signals from the acoustictransducer 1206; and a controller 1211 operatively coupled to theapplanation device 1204 and the signal processor 1208 for controllingthe degree of applanation pressure applied to the artery.

The pressure transducer 1202 is, in the present embodiment, apiezoelectric transducer clement which generates an electrical signal infunctional relationship (e.g., proportional) to the pressure applied toits sensing surface 1212. Similarly, the acoustic transducer 1206comprises a piezoelectric (ceramic) device which is capable of bothgenerating and receiving acoustic waves and/or pulses depending on mode.In the illustrated embodiment, the acoustic transducer 1206 is tuned togenerate ultrasonic frequencies centered at 8 MHz, although other centerfrequencies, with varying bandwidths, may be used. The signalgenerator/receiver 1210 generates electrical signals or pulses which areprovided to the acoustic transducer 1206 and converted into acousticenergy radiated into the blood vessel. This acoustic energy is reflectedby various structures within the artery, including blood flowingtherein, as well as tissue and other bodily components in proximity tothe artery. These acoustic reflections (echoes) are received by theacoustic transducer 1206 and converted into electrical signals which arethen converted by the signal generator/receiver 1210 to a digital form(using, e.g., an ADC) and sent to the signal processor 1208 foranalysis. Depending on the type of acoustic analysis technique and modeemployed, the signal processor 1208 utilizes its program (eitherembedded or stored in an external storage device) to analyze thereceived signals. For example, if the system is used to measure themaximum blood velocity, then the received echoes are analyzed for, interalia, Doppler frequency shift. Alternatively, if the arterial diameter(area) is measured, then an analysis appropriate to the aforementionedA-mode is employed.

During a calibration “sweep”, the controller 1211 controls theapplanation device to applanate the artery (and interposed tissue)according to a predetermined profile. During this sweep, acousticsignals are transmitted into and received from the artery preferably ina region directly proximate the ongoing applanation of the tissue.Velocity, kinetic energy, and/or arterial diameter data is extractedand/or derived from the received echoes and recorded as a function ofthe applanation pressure for the selected portion(s) of the cardiaccycle. The signal processor 1208 and associated algorithms then identifyone or more markers, and determine the desired applied pressure at whichcontinuous monitoring is to occur based on the measured markers. Forexample, if the peak in maximum velocity shown in FIG. 4b were selectedas the marker, the algorithms would identify this peak and identify thepressure data corresponding to this peak. During subsequent bloodpressure monitoring, the controller 1211 would servo the position of theapplanation device 1204 (in the present embodiment, the pressuretransducer 1202) so as to maintain the target pressure, or any othervalue selected by the programmer/user. Subsequent changes in themeasured parameter (e.g., total blood flow kinetic energy) are used toidentify changes in the actual blood pressure within the artery, therebyobviating the need for a continuing series of calibration sweeps.

Optionally, the apparatus 1200 is also configured to measure thetransfer function of the tissue and other bodily components interposedbetween the signal source and the sensor. As described with respect toFIG. 7 above, there is an incomplete or fractional transfer of energybetween the blood within the artery and the pressure sensor. To addressthis issue, the apparatus 1200 of FIG. 12 includes a transfer functionalgorithm (not shown) which utilizes data obtained from A-mode analysisor other techniques relating to the relative compression of the arterialdiameter and the proximate body components when applanated. Hence,during a calibration sweep, the apparatus 1200 stores A-mode or othercomparable data which is used by the transfer function algorithm todetermine the relative compression of the artery and components as afunction of varying applanation pressure. The transfer function (e.g.,change in arterial diameter as a function of applanation pressure) isgenerated by the algorithm and stored in any number of different ways,such as a look-up table or a mathematical function. Subsequent to thecalibration sweep, as the apparatus 1200 servoes to the desired appliedpressure derived from the identified marker, a correction is imposed onthe measured pressure based on the transfer function. For example, ifthe system is servoing to a diastolic pressure of 60 mm Hg as measuredby the pressure transducer 1202, the true value of the pressure in theartery will be corrected according to the transfer function to a valuesomewhat higher than 60 mm Hg.

Referring now to FIG. 13, one specific embodiment of the apparatus forassessing hemodynamic parameters of FIG. 12 is described. In theembodiment of FIG. 13, the apparatus 1300 comprises a self-containedunit having, inter alia, a combined pressure transducer 1302 andapplanation device 1304, acoustic transducer 1306, microcontroller 1308with control micro-code (such as a fuzzy logic algorithm), digitalsignal processor (DSP) 1310 with embedded memory 1312 and instructionset (including calibration lookup tables), signal generator and receiverunit 1314, storage device 1316, display device 1318, and power supply1320. In this embodiment, the microcontroller 1308 is used to controlthe operation of the combined pressure transducer 1302/applanationdevice 1304 so that an initial applanation “sweep” is performed.Specifically, the pressure transducer 1302 is placed in communicationwith the skin of the interior of the wrist of the subject, over theradial artery, and fastened in place as illustrated in FIG. 13.Measurement of the non-calibrated blood pressure from the radial arteryis commenced, and shortly thereafter the microcontroller 1308 directsthe applanation mechanism 1304 to press the transducer 1304 against thewrist of the subject with increasing pressure, thereby applanating theunderlying artery. As the artery is applanated, the acoustic transducer1306 is also pressed in communication with the skin over the artery, andthe signal generator 1314 generates a series of acoustic pulses whichare transmitted through the skin into the artery. As applanation of theartery continues, the signal generator/receiver unit 1314 receivesechoes from the blood and other components within the artery via theacoustic transducer 1306, and generates an output signal relating to thereceived echoes. This output signal is processed and then digitized forsubsequent analysis by the DSP or similar processing engine. Similarly,the output signal of the pressure transducer 1302 is digitized and inputto the DSP. The digitized signals are then analyzed using the embeddedprogram within the DSP, which is a machine code representation of thecomputer program described subsequently herein. The output of thedigital signal processor is a corrected pressure waveform which is thensupplied to the display device 1318 (whether in digital or analog form,depending on the type of device used) for display to the user.Optionally, the output of the DSP may be stored in one or more storagelocations within the storage device 1316, and/or output to an externaldevice.

It is noted that the apparatus 1200, 1300 described herein may beconstructed in a variety of different configurations, using a variety ofdifferent components, and measuring a variety of different hemodynamicparameters. Exemplary control, signal generation/processing, andapplanation mechanisms and circuitry are described in Applicant'sco-pending U.S. patent application, Ser. No. 09/342,549, entitled“Method And Apparatus For The Noninvasive Determination Of ArterialBlood Pressure,” previously incorporated herein.

Computer Program and Related Apparatus

A computer program for implementing the aforementioned methods ofhemodynamic assessment, modeling, and calibration is now described. Inone exemplary embodiment, the computer program comprises an object(“machine”) code representation of a C⁺⁺ source code listingimplementing the methodology of FIGS. 3-3e, 9, and 10, eitherindividually or in combination thereof. While C⁺⁺ language is used forthe present embodiment, it will be appreciated that other programminglanguages may be used, including for example VisualBasic™, Fortran, andC⁺. The object code representation of the source code listing iscompiled and disposed on a media storage device of the type well knownin the computer arts, as illustrated in FIGS. 14a-b. Such media storagedevices can include, without limitation, optical discs, CD ROMs,magnetic floppy disks or “hard” drives, tape drives, or even magneticbubble memory. The computer program further comprises a graphical userinterface (GUI) of the type well known in the programming arts, which isoperatively coupled to the display and input device of the host computeror apparatus on which the program is run (described below with respectto FIG. 15).

In terms of general structure, the program is in one embodimentcomprised of a series of subroutines or algorithms for implementing thehemodynamic assessment, modeling, and calibration methodology describedherein based on measured parametric data provided to the host computer.In a second embodiment, the computer program comprises an assemblylanguage/micro-coded instruction set disposed within the embeddedstorage device, i.e. program memory, of a digital signal processor (DSP)or microprocessor associated with the foregoing hemodynamic measurementapparatus of FIG. 12 or 13.

Referring now to FIG. 15, one embodiment of an apparatus capable ofanalyzing parametric data and generating calibrated values ofhemodynamic parameters as disclosed herein is described. The computingdevice 1500 comprises a motherboard 1501 having a central processingunit (CPU) 1502, random access memory (RAM) 1504, and memory controller(such as a direct memory access controller) 1505. A storage device 1506(such as a hard disk drive or CD-ROM), input device 1507 (such as akeyboard or mouse), and display device 1508 (such as a CRT, plasma, orTFT display), as well as buses necessary to support the operation of thehost and peripheral components, are also provided. A serial or parallelI/O port 1511 is also included for the transfer of data and/or controlsignals to and from the apparatus 1500.

The aforementioned computer program useful for assessing hemodynamicparameters is stored in the form of a machine-readable object coderepresentation in the RAM 1504 and/or storage device 1506 for use by theCPU 1502 during parametric assessment. The user (not shown) assesses thehemodynamic parameters of interest by selecting one or more functionalmodes for the computer program and associated measuring equipment viathe program displays and the input device 1507 during system operation.Specifically, in the case of arterial blood pressure measurement, theuser places the necessary parametric sensors on the selected bloodvessel of the subject, and configures the computer program to acceptdata output by the sensors either continuously or at a predeterminedinterval. The computer program performs the previously describedanalysis if the signals provided to the apparatus 1500, and generates acalibrated signal to be displayed on a display device, or on the systemsown display device. A look-up table or similar mechanism is storedwithin the computer memory or storage device to facilitate calibration,as previously described with respect to FIG. 12, Such calibratedmeasurements generated by the program are also optionally stored in thestorage device 1506 for later retrieval, or output to an external devicesuch as a printer, data storage unit, other peripheral component via aserial or parallel port 1512 if desired. Furthermore, the apparatus 1500may be networked to another computing device or database via networkinterface card (NIC) or similar interface (not shown) whereby the datagenerated by the apparatus 1600 may be remotely analyzed or stored.Transmission to such remote devices may be accomplished using a varietyof well understood methods, such as by local area network (LAN),intranet, Internet, fiber-optic systems, or radio frequency (wireless)devices.

In yet another embodiment, the apparatus comprises a personal computingdevice (such as a personal digital assistant, or PDA), which is adaptedto receive input data from the pressure and acoustic sensors and analyzethe data to produce a corrected measurement of blood pressure. It willalso be recognized that other portable devices, such as laptopcomputers, calculators, and personal organizers, may be configured torun the computer program of the present invention. Furthermore, avariety of different methods of transmitting the input sensor data tothese device may be used, including networked computers, or evenwireless data links.

Method of Providing Treatment

Referring now to FIG. 16, a method of providing treatment to a subjectusing the aforementioned method of assessing hemodynamic parameters isdisclosed. As illustrated in FIG. 16, the first step 1602 of the method1600 comprises monitoring an non-calibrated hemodynamic parameternon-invasively. In the case of blood pressure, an exemplary pressuretransducer applied to the radial artery is used as described withrespect to FIG. 3a herein.

Next, in step 1604 of FIG. 16, a stress is induced on the blood vesselwhich alters its hemodynamic properties (at least locally), therebyinducing changes in other parameters associated with the vessel orcirculatory system as a whole. As previously discussed with respect toFIG. 3b herein, this stress comprises in one embodiment applanating orvariably compressing the blood vessel as a function of time, therebyinducing changes in, inter alia, the volumetric flow (Q), velocity (v),and kinetic energy (KE) of the blood in the region of the applanation.It is again noted, however, that other stressors may conceivably beapplied which may affect similar or other hemodynamic properties.

Next, in step 1606, a second parameter associated with the blood vesselis measured in order to facilitate derivation of a calibration functionin step 1608 below. As discussed with respect to FIG. 3c herein, thesecond parameter in one embodiment comprises total blood flow kineticenergy or maximum blood velocity since these parameter exhibits certaineasily identified “artifacts” as a function of the application of thestressor in step 1604. Other parameters which exhibit the same or otherartifacts may be used to derive the calibration function however.

In step 1608 of FIG. 16, a calibration metric or function is nextderived based on the parametric information derived in step 1606.Specifically, one or more artifacts or markers are identified within theparametric data, these artifacts indicating when certain relationshipsbetween the actual and measured values of the first parameter of step1602 above exist. As discussed with reference to FIGS. 3-3e herein, oneembodiment of the process of deriving a calibration function comprises(i) measuring kinetic energy or maximum blood velocity profilesproximate to the area of the applied stressor (applanation), andidentifying regions of increasing velocity or kinetic energy withinthese profiles as a function of the applanation (correlated topercentage reduction of cross-sectional area of the blood vessel); and(ii) measuring a transfer function for the tissue and other bodilycomponents in the region of pressure measurement.

In step 1610 of the method of FIG. 16, the calibration function derivedin step 1608 is applied to the measurement of the first parameter ofstep 1602 to generate a corrected or calibrated measurement. Note thatif the first parameter is measured continuously (or periodically) as afunction of time, the correction function of step 1608 may becontinuously or periodically applied as appropriate, or alternativelythe second hemodynamic parameter may be monitored (such as duringpressure servoing as previously described) to indicate changes in thecalibration function.

Lastly, in step 1612, the calibrated measurement of the first parameteris used as the basis for providing treatment to the subject. Forexample, in the case of blood pressure measurements, the calibratedsystolic and diastolic blood pressure values are generated and displayedor otherwise provided to the health care provider in real time, such asduring surgery. Alternatively, such calibrated measurements may becollected over an extended period of time and analyzed for long termtrends in the condition or response of the circulatory system of thesubject.

While the above detailed description has shown, described, and pointedout novel features of the invention as applied to various embodiments,it will be understood that various omissions, substitutions, and changesin the form and details of the device or process illustrated may be madeby those skilled in the art without departing from the spirit of theinvention. The foregoing description is of the best mode presentlycontemplated of carrying out the invention. This description is in noway meant to be limiting, but rather should be taken as illustrative ofthe general principles of the invention. The scope of the inventionshould be determined with reference to the claims.

What is claimed is:
 1. A method of assessing hemodynamic propertieswithin the circulatory system of a living subject, comprising: measuringa first parameter from a blood vessel of said subject; compressing saidblood vessel, and measuring a second parameter from said blood vesselduring at least a portion of said act of compressing; identifying atleast one artifact within said second parameter; generating acalibration function based at least in part on said at least oneartifact; and calibrating the first parameter using at least saidcalibration function.
 2. The method of claim 1, wherein the act ofmeasuring a second parameter comprises; transmitting an acoustic waveinto said blood vessel; receiving at least one echo of said acousticwave; and analyzing said at least one echo to derive an estimate of saidsecond parameter.
 3. The method of claim 2, wherein said secondparameter comprises blood velocity, and the act of identifying at leastone artifact comprises identifying a predetermined variation in maximumblood velocity.
 4. The method of claim 2, wherein said second parameteris related to blood velocity, and the act of identifying at least oneartifact comprises identifying a maximum value within said secondparameter.
 5. The method of claim 1, wherein said act of compressingcomprises compressing said blood vessel according to an applanationprofile.
 6. The method of claim 5, wherein said profile comprisesmodulating the position of said at least one transducer as a function oftime.
 7. The method of claim 1, wherein said first parameter comprisespressure, and said second parameter is related to blood flow velocity.8. A method of assessing hemodynamic properties within the circulatorysystem of a living subject, comprising: measuring a pressure waveformfrom a blood vessel of said subject using at least one transducer;compressing said blood vessel using said at least one transducer, andmeasuring a parameter from said blood vessel during at least a portionof said act of compressing; identifying at least one artifact withinsaid parameter; generating a calibration function based at least in parton said at least one artifact; and calibrating the pressure waveformusing at least said calibration function.
 9. The method of claim 8,wherein said parameter comprises blood velocity, and the act ofidentifying at least one artifact comprises identifying a predeterminedvariation in maximum blood velocity.
 10. The method of claim 8, whereinsaid parameter comprises blood flow kinetic energy.
 11. The method ofclaim 8, wherein said act of compressing comprises compressing saidblood vessel according to an applanation profile.
 12. The method ofclaim 11, wherein said profile comprises modulating the position of saidat least one transducer as a function of time.
 13. A method of assessinghemodynamic properties within the circulatory system of a livingsubject, comprising: measuring a first parameter from a blood vessel ofsaid subject; compressing said blood vessel, and measuring from saidvessel the kinetic energy of blood flowing therein during at least aportion of said act of compressing; deriving a calibration functionbased at least in part on said kinetic energy; and calibrating the firstparameter using said calibration function.
 14. The method of claim 12,wherein the act of measuring the kinetic energy comprises: measuring theDoppler shift associated with the blood flowing within the blood vessel;and calculating the mean kinetic energy of said blood based on saidDoppler shift.
 15. A method of assessing hemodynamic properties withinthe circulatory system of a living subject, comprising: measuring afirst parameter from a blood vessel of said subject; compressing saidblood vessel. and measuring the diameter of said blood vessel during atleast a portion of said act of compressing; deriving a calibrationfunction based at least in part on said diameter; and calibrating thefirst parameter using said calibration function.
 16. The method of claim15, wherein the act of measuring said diameter comprises: transmittingan acoustic wave into said blood vessel; receiving at least one echofrom said acoustic wave; and analyzing said at least one echo todetermine the arterial diameter.
 17. The method of claim 16, whereinsaid acoustic wave comprises an A-mode acoustic transmission.
 18. Themethod of claim 16, wherein said acoustic wave comprises an M-mode(motion-mode) acoustic transmission.
 19. The method of claim 16, whereinsaid acoustic wave comprises a B-mode (brightness mode) acoustictransmission.
 20. The method of claim 15, wherein said act of deriving acalibration function comprises identifying at least one artifact duringsaid act of measuring said diameter, and generating said calibrationfunction based at least in part on said at least one artifact.
 21. Themethod of claim 20, wherein said act of identifying at least oneartifact comprises identifying the level of compression at which thecross-sectional area of said blood vessel is maximized.
 22. A method ofassessing hemodynamic properties within the circulatory system of aliving subject, comprising: measuring a first parameter from a bloodvessel of said subject; compressing said blood vessel, and measuring thevolumetric flow rate associated with at least a portion of said bloodvessel during at least a portion of said act of compressing; deriving acalibration function based at least in part on said volumetric flowrate; and calibrating the first parameter using said calibrationfunction.
 23. The method of claim 22, wherein said first parametercomprises pressure, and said act of deriving a calibration functioncomprises identifying at least one artifact within acompression/volumetric flow rate profile, and deriving a scalingfunction based at least in part on the pressure measured at the pointwhere said at least one artifact occurs.
 24. A method of assessinghemodynamic properties within the circulatory system of a livingsubject, comprising: measuring a first parameter from a blood vessel ofsaid subject; compressing said blood vessel, and measuring a secondparameter from said blood vessel during at least a portion of said actof compressing; measuring the actual compression of said blood vesselduring at least a portion of said act of compressing; generating atransfer function indicative of said actual compression; deriving acalibration function based at least in part on said second parameter andsaid transfer function; and calibrating the first parameter using saidcalibration function.
 25. The method of claim 24, wherein said act ofgenerating a transfer function comprises generating at least onefunction relating level of compression and flow area.
 26. The method ofclaim 24, wherein said act of generating a transfer function comprisesgenerating separate functions for systolic and diastolic conditions. 27.A method of measuring the pressure within the circulatory system of aliving subject, comprising: measuring a non-calibrated pressure from ablood vessel of said subject; measuring a second parameter from saidblood vessel; creating a stress on said blood vessel during at least aportion of the act of measuring said second parameter, said stressproducing a known effect on said second parameter, the application ofsaid stress being selectively controlled so as to maintain apredetermined condition within said blood vessel; identifying at leastone marker within said measurement of said second parameter, said atleast one marker being related to said known effect; deriving acalibration function based at least in part on said at least one marker;and correcting said non-calibrated pressure using the derivedcalibration function.
 28. The method of claim 27, wherein the act ofmeasuring a second parameter comprises measuring said second parameterusing an acoustic wave, and the act of creating a stress comprisesvariably compressing said blood vessel as a function of time.
 29. Themethod of claim 28, wherein said predetermined condition comprisesmaintaining the pressure across the wall of the blood vessel within apredetermined range.
 30. The method of claim 27, wherein saidpredetermined condition comprises maintaining the pressure across thewall of the blood vessel within a predetermined range.
 31. A method ofmeasuring the pressure within the circulatory system of a livingsubject, comprising: measuring a non-calibrated pressure from a bloodvessel of said subject; measuring the velocity of blood flowing withinsaid blood vessel using an acoustic wave and by forming a time-frequencyrepresentation from echoes of said acoustic wave; variably compressingsaid blood vessel as a function of time during at least a portion of theact of measuring said velocity, said compressing producing a knowneffect on said velocity; identifying at least one marker within saidmeasurement of said velocity, said at least one marker being related tosaid known effect; deriving a calibration function based at least inpart on said at least one marker; and correcting said non-calibratedpressure using the derived calibration function.
 32. A method ofmeasuring the pressure within the circulatory system of a livingsubject, comprising: measuring a non-calibrated pressure waveform from ablood vessel of said subject; determining the velocity of blood flowingwithin said blood vessel at least in part by forming a time-frequencyrepresentation from echoes of acoustic waves transmitted into said bloodvessel; variably compressing said blood vessel as a function of timeduring at least a portion of the act of determining said velocity, saidcompressing producing a known effect on said velocity; identifying atleast one marker within said determination of said velocity, said atleast one marker being related to said known effect; deriving acalibration function based at least in part on said at least one marker;and correcting said non-calibrated pressure waveform using the derivedcalibration function.
 33. A method of measuring the pressure within thecirculatory system of a living subject, comprising: measuring anon-calibrated pressure from a blood vessel of said subject; measuringthe velocity of blood flowing within said blood vessel using an acousticwave; forming a time-frequency representation from echoes of saidacoustic wave; variably compressing said blood vessel as a function oftime during at least a portion of the act of measuring said velocity,said compressing producing a known effect on said velocity; identifyingat least one marker within said measurement of said velocity, said atleast one marker being related to said known effect, said act ofidentifying being based at least in part on said time-frequencyrepresentation; deriving a calibration function based at least in parton said at least one marker; and correcting said non-calibrated pressureusing the derived calibration function.
 34. An apparatus for measuringhemodynamic properties within the radial artery of a human subjectcomprising: a pressure transducer adapted to be located proximate saidartery and measure at least a pressure waveform therefrom; an acoustictransducer adapted to measure at least one hemodynamic parameterassociated with said artery; an applanation device adapted to compressat least a portion of said artery while measuring said at least onehemodynamic parameter; and a signal processor operatively connected tosaid pressure and acoustic transducers and configured to generate acalibration function based at least in part on a signal produced by saidacoustic transducer, and apply said calibration function to a signalproduced by said pressure transducer.
 35. The apparatus of claim 34,wherein said acoustic transducer transmits at least one acousticemission into said artery and receives at least one echo therefrom, saidat least one echo being used by said signal processor to generate arelationship between said at least one hemodynamic parameter and thepressure applied by said applanation device.
 36. The apparatus of claim34, wherein said applanation device comprises said pressure transducer.37. The apparatus of claim 35, wherein said at least one acousticemission comprises an acoustic pulse.
 38. The apparatus of claim 35,wherein said at least one acoustic emission comprises a substantiallycontinuous acoustic wave.
 39. The apparatus of claim 34, wherein said atleast one hemodynamic parameter comprises the maximum velocity of bloodflowing within said artery.
 40. The apparatus of claim 34, wherein saidat least one hemodynamic parameter comprises the amplitude-weighted meanvelocity of blood flowing within said artery.
 41. The apparatus of claim34, wherein said at least one hemodynamic parameter is related to thekinetic energy of blood flowing within said artery.
 42. An informationstorage device, comprising; a data storage medium; a plurality of datastored on said medium, said plurality of data comprising a computerprogram adapted to run on a data processor, said computer program beingconfigured for determining a calibrated value of a first hemodynamicparameter of a living subject based on first, second, and third inputdata, said first input data being representative of a non-calibratedvalue of said first hemodynamic parameter, said second input data beingrepresentative of a second hemodynamic parameter, said third input datarelating to the error associated with said first input data, the act ofdetermining said calibrated value of said first hemodynamic parametercomprising: analyzing said second input data to identify at least onefunctional relationship therein; identifying at least one marker withinsaid at least one functional relationship; determining a calibrationfunction based at least in part on said at least one marker and saidthird data; and applying said calibration function to said first inputdata to determine said calibrated value.
 43. An information storagedevice, comprising; a data storage medium; a plurality of data stored onsaid medium, said plurality of data comprising a computer programadapted to run on a data processor, said computer program beingconfigured for determining a calibrated value of a first hemodynamicparameter of a living subject based on first input data and second inputdata, said first input data being representative of a non-calibratedvalue of said first hemodynamic parameter, said second input data beingrepresentative of a second hemodynamic parameter, the act of determiningsaid calibrated value of said first hemodynamic parameter comprising:forming a time-frequency distribution from said second input data;analyzing at least said time-frequency distribution to identify at leastone artifact therein; utilizing said at least one artifact to derive acalibration function; and applying said calibration function to saidfirst input data to determine said calibrated value.